This invention relates generally to improving the quality of X-ray images, such as computed tomographic (CT) images, and more particularly to an improved system for and method of using readings of X-ray exposure taken during a CT scan to normalize image data derived from that scan.
CT scanners of the third generation type include an X-ray source and X-ray detector system secured respectively on diametrically opposite sides of an annular-shaped disk. The latter is rotatably mounted within a gantry support so that during a scan the disk continuously rotates about a rotation axis while X-rays pass from the source through an object positioned within the opening of the disk to the detector system.
The detector system typically includes an array of detectors disposed in the shape of an arc of a circle having a center of curvature at the point, referred to as the xe2x80x9cfocal spot,xe2x80x9d where the radiation emanates from the X-ray source. The X-rays that are detected by a single detector at a measuring instant during a scan is considered a xe2x80x9cray.xe2x80x9d The ray is partially attenuated by all the mass in its path so as to generate a single intensity measurement as a function of the attenuation, and thus the density of the mass in that path. The X-ray source and array of detectors are all positioned so that the X-ray paths between the source and each detector define a xe2x80x9cray pathxe2x80x9d. Because the ray paths originate from substantially a point source and extend at different angles to the detectors, the ray paths resemble a fan in the case of a single row of detectors, or a cone in the case of multiple rows of detectors. Projections or views, i.e., the X-ray intensity measurements, are typically done at each of a plurality of angular positions of the disk. The data can be acquired with the scanned object still, as for example, with a constant axis scan, or while the object and rotating disk are moved relative to one another in the direction of the Z-axis (the axis of rotation of the disk carrying the x-ray source and detector array) while the source and detector array rotate about the object, as for example, with a helical scan. The detector array can be symmetrical or asymmetrical about a ray passing from the source through the isocenter of the machine, i.e., the center of rotation of the X-ray source and detectors.
An image reconstructed from data acquired at all of the projection angles during the scan will be a slice made up of data acquired along the scanning section through the object being scanned. In the case of a constant axis scan the slice will be a plane passing through the Z-axis, while in the case of a helical scan the slice will be a volume portion of the object through which rays pass as the object is moved in the Z-axis direction. In order to xe2x80x9creconstructxe2x80x9d a density image of the section or xe2x80x9cslicexe2x80x9d of the object within the xe2x80x9cfield of viewxe2x80x9d in the defined scanning plane, the image is typically reconstructed in a pixel array, wherein each pixel in the array is attributed a value representative of the attenuation of all of the rays that pass through its corresponding position in the scanning plane during a scan. As the source and detectors rotate around the object, rays penetrate the object from different directions, or projection angles, passing through different combinations of pixel locations. The density distribution of the object in the slice section is mathematically generated from these measurements, and the brightness value of each pixel is set to represent that distribution. The result is an array of pixels of differing values which represents a density image of the slice plane.
As described in U.S. Pat. No. 5,680,427 (hereinafter the ""427 Patent), issued to John M. Dobbs et al. and assigned to the present assignee, in order to produce a good quality image, the CT scanner designer works hard to minimize sources of error. Accordingly, steps are usually taken to provide for correction of errors either through design or calibration. For example, at zero X-ray levels it is important to minimize and stabilize signal offsets so that any measurement will contain a known constant offset for which corrections can be made. In addition, X-rays are provided at full scale and measurements are taken so as to generate xe2x80x9cairxe2x80x9d data with no absorbent material in the path of the X-rays so as to minimize errors due to drift in gain and measurement uncertainty at full scale. Two points of reference are thus provided between which data is corrected. In between these two points representing zero and full scale there is a curve which represents the relationship between X-ray levels and data values. The non-linear relationship between X-ray levels and data values results because the electrical signal varies in a non-linear manner with signal strength. Accordingly, materials of known absorption values (e.g., water, polyethylene, polyvinyl chloride, etc.) of predetermined thicknesses are placed within the path of the fan beam and data are generated in order to calibrate the system. The data will represent points on the curve. Using these known materials allows for the determination of the correct dosage level for a particular scan, and detector efficiency. A best fit polynomial can be easily determined using known techniques so that a look up table can be generated and stored.
Within the context of insuring good tomographic images, it is also important that the data represent identical detection for all of the detectors for any given number of photons. If one datum, representative of a number of photons received during the measurement period from one detector is different from the data received from all of the other channels for the same measurement, the result will be an artifact in the reconstructed image. Thus, steps have been taken in the past to calibrate the offset and gain of each data channel so that errors attributed to these two factors are minimized.
Additional errors are attributable to the source of X-rays. Even though an X-ray tube is set to provide a constant X-ray flux output, the number of photons striking the detectors within a prescribed period of time can vary from detector to detector. As mentioned above, it is also known that each photon contributes to noise. Thus, the fewer number of photons detected, the poorer the signal-to-noise ratio (S/N).
In addition, the X-ray source may fluctuate during the scan, particularly as it reaches the end of its xe2x80x9clifexe2x80x9d, producing fluctuating intensities of the X-rays; and in at least one case even though the X-ray source is set to provide a given number of photons for each view, the signal can be degraded.
Typically, in order to account for fluctuating intensities of the X-rays, the CT data is normalized. More particularly, one of the first steps of a CT reconstruction algorithm is to calculate how much the X-rays are attenuated by the scanned object. This is accomplished by dividing the signal measured without an object in the X-ray path (called air scan) by the signal measured while scanning the object. A monitor (also called reference) detector is usually located so that the X-rays always pass from the source to the monitor detector without being attenuated by an object. The monitor data is used to account for fluctuations in the incident X-ray intensity during the air scan and the object scan. If a monitor detector is obstructed, the monitor data no longer reflects the incident X-ray intensity, which can lead to image artifacts.
Thus, it is known to utilize a monitor detector system to monitor the level of x-rays emitted by the X-ray source during a scan, and to use the monitor values to normalize the CT data taken at that time. The monitor detector system can be separate from the arcuate image detector array and positioned so that it remains clear of the object being scanned. See, for example, the ""427 Patent. In some cases, where it may be desirable to eliminate a separate monitor detector system, because of the added costs it provides, detectors of the detector array can be used to acquire monitored data. For example, U.S. Pat. No. 4,769,827 issued Sep. 6, 1988 to Uno, et al., (the ""827 Patentxe2x80x9d) provides a pair of reference detectors respectively at opposite ends of the arcuate image detector array to provide reference signals by which the data signals provided during a view can be compared, i.e., normalized. In one commercially available CT scan system, two detectors at the asymmetric end of the detector array are used as reference detectors, and obviously, the two end detectors of the symmetric end could be used. The problem with using detectors of the arcuate image detector array can be that during a scan, the rays from the X-ray source to the reference detectors is partially attenuated by the support table and object being scanned for certain positions of the gantry. As a result, monitor readings for those obstructed views are not valid for reference correction. In one prior art approach the reconstruction algorithm which uses the scanned data to reconstruct an image of the slice taken, uses an airtable-based threshold to replace low monitor readings. Thresholding obstructed monitor readings still leads to image artifacts because the monitor value is clipped at a value which is below the typical range for the scan.
An improved method of and system for correcting for X-ray fluctuations is provided. The adaptive monitor correction method and system uses all monitor readings for a scan to find views where monitor detectors are obstructed. Reference readings for obstructed views are replaced with one or more values calculated based on the remaining valid monitor readings. The values for the obstructed views are preferably statistically determined.
In one preferred embodiment a histogram of input monitor values and low-pass filtered monitor data are used in finding obstructed monitor views. Parameters controlling the algorithm performance were chosen after testing on several sets of raw data collected on the scanner. Monitor-related image artifacts are reduced by the adaptive monitor correction algorithm described hereinafter.